KR20250052496A - Blood pump - Google Patents

Abstract

The present invention relates to a control device (100) for controlling blood flow (Q _pump (t)) of an intravascular blood pump (50) for percutaneous insertion into a blood vessel of a patient, wherein the blood pump (50) comprises a pump unit (52), and a drive unit (51) for driving the pump unit configured to deliver blood from a blood flow inlet (54) toward a blood flow outlet (56), wherein the control device (100) is configured to operate the blood pump (50) in a selectable zero flow control mode - a blood flow command signal (Q _pump ^set (t)) is selected - and the control device comprises a first controller (401) and a second controller (402), wherein the first controller (401) is configured to control the blood flow (Q pump (t)) by regulating a speed command signal (n _pump set (t)) to the drive unit (51), and the second controller (402) is configured to control the blood flow (Q _pump (t)) by regulating a speed command signal (n pump ^set (t)) to the drive unit (51). The controller (402) is configured to control the driving speed (n _pump (t)) of the drive unit (51). The present invention also relates to a corresponding method for controlling blood flow (Q _pump (t)) of an intravascular blood pump (50), as well as a control device (100) configured to execute the method. Finally, the present invention relates to a system comprising an intravascular blood pump (50) for assisting the heart and a control device (100) for controlling the blood pump (50).

Description

BLOOD PUMP

The present invention relates to blood pumps, and particularly to intravascular blood pumps for percutaneous insertion into a blood vessel of a patient. In particular, the present invention relates to a specific method for controlling a percutaneously implantable blood pump, and to a corresponding control device, as well as to a system comprising the control device and the blood pump. The present invention is configured for and particularly useful for intravascular blood pumps, but is less relevant for larger blood pumps, such as VADs, which are not positioned within a blood vessel or within the heart, but are implanted outside the heart of a patient, for example in the thoracic cavity.

A ventricular assist device (VAD) may be used to support the function of a patient's heart as a left ventricular assist device (LVAD) or a right ventricular assist device (RVAD). A typical VAD is connected to the patient's heart by a suitable conduit and is implanted into the patient's chest cavity outside the heart, whereas an intravascular blood pump for percutaneous insertion typically comprises a catheter and a pump unit, which is inserted into the patient's heart via an intravascular access, for example, through the aorta into the left ventricle. The pump unit is positioned at the distal end of the catheter and comprises a blood flow inlet, a blood flow outlet and a cannula, through which blood flow is generated, for example, by a rotor or impeller of the pump unit. For example, the cannula may extend through the aortic valve with the blood flow inlet positioned at the distal end of the cannula in the left ventricle and the blood flow outlet positioned at the proximal end of the cannula in the aorta. By creating blood flow, the pressure difference between the outlet and inlet is overcome.

First and foremost, an important aspect of intravascular blood pumps (hereinafter also simply referred to as "blood pumps") is to explant the intravascular blood pump from the patient to ensure that natural heart function has been restored. This can be done, for example, by appropriately reducing the blood flow provided by the blood pump so that the blood pump can finally be explanted after it has been determined that the heart has sufficiently recovered. This aspect, i.e., determining the exact timing of explantation, is not as important in larger VADs, which are typically implanted in the patient's thoracic cavity and are designed for long-term use.

To date, no physical signal is known that adequately represents the state of cardiac recovery while an intravascular blood pump is implanted. While the blood pump is assisting the heart, it is not possible to know the function of the heart in the unassisted state. Furthermore, when the blood pump is switched off, flow reflux through the cannula occurs, making it impossible to know the function of the heart in the unassisted state. Reflux is a problem in particular for endovascular pumps, since the blood pump, more specifically the cannula of the pump, extends through a heart valve, e.g. the aortic valve, creating an open path through the heart valve, i.e. backflow into the heart when the blood pump is not driven. This issue does not occur with extravascular devices, such as VADs, which are placed outside the heart, e.g. in the thoracic cavity, because they do not extend through the heart valve but bypass it.

In the state of the art, the current pump speed setting is manually reduced gradually, for example by one level, by the physician based on his/her professional experience. After the reduction of the pump speed, the mean aortic pressure is monitored. In some institutions, an ultrasound-based assessment of the left ventricular volume and continuous cardiac output measurements are performed. If the mean aortic pressure remains stable, it is assumed that the heart can take over the work performed by the blood pump. However, if the mean aortic pressure drops, it is assumed that the heart still needs more assistance and the pump speed needs to be increased again. In addition, before explantation of the blood pump, the so-called on/off approach is applied. In this case, the pump speed is significantly reduced, for example, for several hours, and the physiological condition of the patient and in particular the ventricular dilatation are observed, for example, based on echocardiography (ECHO) measurements and/or ventriculography. ECHO can provide information about the heart, such as size and shape, size quantification, pumping capacity, and allow calculation of cardiac output, ejection fraction, and diastolic function. Cardiac ventriculography involves injecting contrast media into the ventricles of the heart to measure the volume of blood being pumped. The measurements obtained by cardiac ventriculography are ejection fraction, stroke volume, and cardiac output.

When the blood pump is switched off, and if the heart is still working insufficiently, the ventricle, which is no longer assisted, is greatly dilated, so that sufficient blood volume is not ejected from the ventricle during systole, leading to an increase in left ventricular end-diastolic volume and pressure. In other words, due to the reduced pump speed, the heart is subjected to a load that it can cope with, which is still a severe overload of the heart that has not yet recovered. This can result in several days of treatment interruption.

Therefore, the actual monitoring process prior to explantation of a blood pump is somewhat of a trial and error process, with the pump speed needing to be reduced further if the patient's condition remains stable, and then increased again if it worsens.

It is an object of the present invention to provide an improved control method for an intravascular blood pump and a correspondingly improved control device, as well as a system comprising said control device and an intravascular blood pump, wherein said blood pump can be operated so that a better assessment of the state of cardiac recovery can be ascertained.

This object is achieved by the features of each independent claim. Advantageous embodiments and further improvements are defined in each dependent claim.

For clarity, the following definitions will apply here:

The term "characteristic parameters of the heart" is to be understood as specific values derived from physiological signals which can characterize the state of the heart, for example loading (overloaded or unloaded) and/or physiological conditions (e.g. weak, healthy or recovering).

The "human circulatory system" is the organ system that circulates blood. The essential components of the human circulatory system are the heart, blood, and blood vessels. The circulatory system includes the pulmonary circulation (the "loop" through which blood is oxygenated by the lungs); and the systemic circulation (the "loop" through which blood is oxygenated by the rest of the body).

The improvements disclosed herein relate to a blood pump comprising a settable blood flow level. For example, in the case of a rotary blood pump, the "settable blood flow level" may be a discrete blood flow level within a range defined by a minimum blood flow and a maximum blood flow, or a continuously settable blood flow level.

The basic idea of the control device and the corresponding control method for controlling an intravascular blood pump proposed herein is to provide a mode in which the current blood flow through the blood pump can be kept very low, preferably zero, compared to the blood flow outlet capacity. Preferably, the blood flow through the blood pump is kept between 0 and 1 L/min, more preferably between 0 and 0.5 L/min, between 0 and 0.2 L/min or even between 0 and 0.1 L/min. Most preferably, the blood flow through the blood pump is kept at almost zero flow. In this case, the blood pump is controlled such that neither a positive blood flow nor a negative blood flow is generated. This mode of operation is referred to herein as "zero flow control mode". For example, zero flow can be established or maintained by controlling the current driving speed of the drive unit (e.g., a motor) of the blood pump, and in particular, by compensating for the current pressure difference between the blood flow inlet and the blood flow outlet of the blood pump.

In this context "zero flow" should be understood as zero flow or very low blood flow. Since the purpose of the zero flow mode is to obtain knowledge about the recovery state of the heart, performing very low blood flows may also be appropriate. Low blood flows, for example below 0.1 L/min, below 0.2 L/min or even below 0.5 L/min should be considered as "zero flow" in this context. In no case should "zero flow" be negative. In other words, zero flow does not allow any backflow through the blood pump.

In this context, an intravascular blood pump for percutaneous insertion comprises a catheter and a pump unit, which is inserted via a blood vessel into a patient's heart, for example, through the aorta into the left ventricle. The pump unit comprises a blood flow inlet, a blood flow outlet, and a cannula, through which a drive unit for driving the pump unit generates blood flow. For example, the pump unit may comprise a drive unit, for example, a rotor or impeller driven by a motor, to deliver blood from the blood flow inlet toward the blood flow outlet. For example, the cannula may extend through the aortic valve with the blood flow inlet positioned at a distal end of the cannula within the left ventricle and the blood flow outlet positioned at a proximal end of the cannula within the aorta. The intravascular blood pump can have a maximum outer diameter in the range of about 12 French (F) (about 4 mm) to 21 French (F) (about 7 mm), for example, 12 F (about 4 mm), 18 F (about 6 mm) or 21 F (about 7 mm), which is typically the maximum outer diameter of the pump unit. The catheter can have an outer diameter that is smaller than the outer diameter of the pump unit, for example, 9 F (about 3 mm).

Natural heart function creates a pressure difference between the aorta and the left ventricle, for example. In order to create positive blood flow, the blood pump must overcome this pressure difference. Otherwise, that is, if the pressure created by the blood pump is too low, the pressure difference between the aorta and the left ventricle will cause backflow into the left ventricle.

By applying the zero flow control mode, the blood pump provides no assistance or very low assistance to the heart, and advantageously avoids backflow of blood. That is, the blood pump does not allow backflow of blood. For example, in the case of left ventricular assistance, the blood pump does not allow backflow of blood (black-flow) from the aorta back into the left ventricle during diastole of the heart.

By applying zero flow control mode, the drive unit of the blood pump, e.g. the rotor or impeller, continues to rotate. Therefore, the risk of thrombosis due to continuously moving parts is reduced.

By zero flow mode of operation, the assistance provided to the heart by the blood pump will be set to be essentially zero. "Essentially zero" means that any blood flow still generated must be at least negligible, but in no case negative. That is, the blood pump does not allow backflow through the blood pump.

In the zero flow mode of operation, the complete work of overcoming the pressure difference between the pressure in an auxiliary ventricle, for example the left ventricle, and the pressure in an adjacent blood vessel, for example the aorta, is performed solely by the heart. In this way, the zero flow mode of operation allows monitoring one or more suitable characteristic parameters of the heart that can be interpreted or used as indicators of the state of cardiac recovery.

Preferably, the blood flow of the blood pump is related to the driving speed of the drive unit, e.g. the motor, the current supplied to the drive unit and/or the pressure difference between the outlet and the inlet of the blood pump. These relationships can be stored in a memory, e.g. a look-up table as described in more detail below. In other words, the command signal values can be stored in a memory of the control device or in a memory within the blood pump accessible by the control device.

A first aspect provides a control device for controlling blood flow (Q _pump (t)) of an intravascular blood pump for percutaneous insertion into a blood vessel of a patient. The blood pump comprises a pump unit, and a drive unit for driving the pump unit configured to deliver blood from a blood flow inlet toward a blood flow outlet. The control device is configured to operate the blood pump in a selectable zero flow control mode, wherein a blood flow command signal (Q _pump ^set (t)) is selected. The control device comprises a first controller and a second controller, the first controller being configured to control the blood flow (Q _pump (t)) by regulating a speed command signal (n _pump ^set (t)) to the drive unit, and the second controller being configured to control a driving speed (n _pump (t)) of the drive unit. More specifically, the control device is configured specifically for an intravascular blood pump, or more generally, for a low-inertia device as described in more detail below.

Preferably, the intravascular blood pump comprises a cannula between the blood flow inlet and the blood flow outlet, through which blood flow is generated by the pump unit. In operation, the cannula may extend through the aortic valve, for example, with the blood flow inlet positioned in the left ventricle and the blood flow outlet positioned in the aorta.

For example, the controlled blood flow can be constant. By compensating for the current pressure difference between the blood flow inlet and the blood flow outlet, the actual blood flow through the blood pump results in zero flow. In other words, in the zero flow control mode, the current pressure difference between the blood flow outlet and the blood flow inlet is compensated by controlling the blood flow through controlling the driving speed.

Preferably, the first controller is configured to determine the velocity command signal (n _pump ^set (t)) based on the difference (ΔQ) between the blood flow command signal (Q _pump ^set (t)) and the blood flow (Q _pump (t)). In other words, the first controller is configured to compare the blood flow command signal (Q _pump ^set (t)) with ^the actual blood flow (Q _pump (t)) to determine the velocity command signal (n _pump set (t)).

Preferably, the second controller is configured to control the drive speed (n _pump (t)) by regulating the drive current (I _pump (t)) supplied to the drive unit. For example, the drive unit may comprise a motor, in particular an electric motor, and the regulated drive current may be a motor current supplied to the motor. Thus, in the case of a rotary drive unit, the set drive speed (n _pump ^set (t)) of the drive unit and the command speed signal (n _pump ^set (t)) may be the rotational speed. The motor may be located within the pump unit and may be directly or indirectly coupled to the impeller, for example by a mechanical connection or a magnetic coupling.

Preferably, the first controller and the second controller are part of a cascade control system, wherein in the cascade system, the first controller is an outer controller and the second controller is an inner controller. The outer controller is embedded in the outer control loop and can regulate the generated blood flow by comparing the blood flow command signal with the blood flow generated by the blood pump and by setting a set-point of the inner control loop, i.e., the speed command signal of the blood pump. The inner controller is part of the inner control loop and can control the speed of the blood pump by adjusting the motor current accordingly.

Preferably, the control device is configured to control the blood flow (Q _pump (t)) for a preset zero flow control period.

For example, the preset zero flow control period can be set to last for a portion of one cardiac cycle of the assisted heart. That is, the zero flow control mode is applied only briefly "within a beat". In such a case, the preset zero flow period is preferably small compared to the duration of the cardiac cycle. In this way, information about the recovery state of the heart can be collected without any cardiac overload, since the duration of no assistance to the heart is kept to a minimum.

For example, the preset zero flow control period can be set to last for at least one complete cardiac cycle or a preset number of consecutive complete cardiac cycles.

Preferably, the control device is configured to synchronize the occurrence of at least one characteristic cardiac cycle event with the zero flow control period. For example, the start and/or the end of the zero flow control period is synchronized with the occurrence of at least one characteristic cardiac cycle event. In particular, the start and the end of the zero flow control period can be synchronized with the occurrence of two characteristic cardiac cycle events. In this way, the zero flow control mode can be set to a time interval of the cardiac cycle in which certain characteristic parameters of the heart can provide useful specific information directly or indirectly indicating the state of cardiac recovery.

For example, a characteristic cardiac cycle event may be the opening of the aortic valve or the closure of the aortic valve. For example, the control device may be configured to detect the opening of the aortic valve by one of the following: the presence of a balance between left ventricular pressure and aortic pressure, and the occurrence of an R wave in an electrocardiogram (ECG) signal from a patient having an auxiliary heart.

Additionally, characteristic cardiac cycle events may be opening of the left atrioventricular valve, closure of the left atrioventricular valve, or development of end diastolic left ventricular pressure.

Preferably, the control device is configured to monitor the values of one or more characteristic cardiac parameters. That is, the control device may be configured to monitor, in the zero flow control mode, the one or more characteristic cardiac parameters whenever the zero flow control mode is applied.

Preferably, the control device is configured to operate the intravascular blood pump in a zero flow control mode periodically or randomly. The periodic or random application of the zero flow control mode can be performed over a preset period of time, for example, from a portion of the cardiac cycle to several days or less.

Preferably, the control device is configured to identify a trend in the values of one or more of the monitored characteristic cardiac parameters. The trend of the one or more monitored characteristic cardiac parameters can thereby be used as an indicator of the state of recovery of the heart or the state of cardiac recovery, and therefore, i.e. the trend can be helpful in determining whether recovery is progressing. The trend can be displayed to a physician via a user interface of the control device so that the physician can make a decision about the state of cardiac recovery.

For example, at least one characteristic parameter of the heart may be an arterial blood pressure, which is measured whenever a zero-flow operating mode is established. By applying the zero-flow control mode, the arterial blood pressure may drop. A pressure drop that reaches a critical value or exhibits a critical decrease indicates that the heart has not recovered and therefore it is not possible to explant the blood pump. In another example, the arterial blood pressure may remain stable or exhibit only a minor pressure drop during the zero-flow control mode. In such a case, it may be assumed that the heart has recovered sufficiently and that the blood pump can be explanted.

Preferably, the at least one characteristic cardiac parameter is at least one of arterial pressure pulsatility (AOP| _max - AOP| _min ), mean arterial pressure, cardiac contractility (dLVP(t)/dt| _max ), cardiac relaxation (dLVP(t)/dt| _min ), and heart rate (HR).

The control device may be configured to measure the blood flow (Q _pump (t)) by the sensors, or to calculate or estimate the blood flow (Q _pump (t)). For example, the pressure difference between the blood flow outlet and the blood flow inlet may be determined by respective pressure sensors positioned at the inlet and outlet of the blood pump, i.e., a pressure sensor capturing the after-load of the blood pump and a pressure sensor capturing the pre-load of the blood pump. The blood pump may alternatively or additionally include only one sensor configured to directly measure the pressure difference. Also alternatively, the pressure difference between the blood flow outlet and the blood flow inlet may be estimated, measured or calculated.

Rather than measuring the blood flow (Q _pump (t)), the blood flow (Q pump (t)) may be determined using a look-up table that can describe the relationship between the blood flow, the drive speed, the pressure difference between the blood flow outlet and the blood flow _inlet , and at least one of the drive current supplied to the drive unit. The look-up table may include a set of characteristic curves describing each of the relationships, for example, a set of curves for each of a particular pump speed. It will be appreciated that other suitable look-up tables may be used and that the values in the look-up table may be given in various units.

Data for use in the lookup table, e.g. motor current and blood flow, can be recorded in a test bench assembly by driving the blood pump in the fluid at a given motor speed and at a defined pump load (pressure difference between the inlet and outlet) while recording the flow produced by the pump. The pump load can be increased over time, e.g. from zero load (no pressure difference between the blood flow inlet and the blood flow outlet, i.e. maximum flow) to maximum load (no pump function, i.e. no flow), while the motor current and blood flow are recorded. Such a lookup table can be generated for several different motor speeds. Using such a lookup table to determine blood flow (Q _pump (t)) can provide an advantageous way to determine blood flow (Q _pump (t)) during operation of the blood pump, particularly as compared to measuring or calculating blood flow (Q _pump (t)). With the lookup table, the blood flow (Q _pump (t)) can be determined based on readily available operating parameters of the blood pump alone. Thus, no sensors for detecting parameters of the patient, for example, a pressure sensor or a flow sensor for detecting a pressure difference inside the patient's blood vessels, are required. Furthermore, reading the value for the blood flow (Q _pump (t)) from the lookup table does not require computationally intensive calculations.

However, if one or more suitable characteristic cardiac parameters are monitored by only applying the zero flow control mode within a single beat, the heart may not be sufficiently adapted to the lost assistance of the blood pump. Therefore, the monitored characteristic cardiac parameters may still not sufficiently represent the exact state of cardiac recovery, e.g. the actual pumping capacity of the heart. Therefore, it is possible to repeat the zero flow control mode within a single beat in several consecutive cardiac cycles.

Thus, the preset zero flow period can be set to last for at least one complete cardiac cycle or a preset number of consecutive complete cardiac cycles. For example, the preset zero flow period can be set to be a portion of a cardiac cycle of several hours or less. In this way, the heart can be sufficiently adapted to the condition of zero assistance by the blood pump, so that the actual state of cardiac recovery can be better confirmed.

Additionally, it is possible to couple zero flow within a single beat over a complete cardiac cycle. For example, the zero flow control mode can initially be applied for a relatively short period of time, for example, a portion of a cardiac cycle from 1 to 300 consecutive cardiac cycles. After verification of natural cardiac function and sufficient recovery, the zero flow control mode can be applied over a longer period, for example, several minutes or hours to several days over a complete cardiac cycle.

A second aspect provides a system comprising a control device according to the first aspect and an intravascular blood pump for assisting the heart.

Preferably, the blood pump is catheter-based, i.e. the blood pump preferably comprises a catheter and a pump unit, preferably with the pump unit positioned at the distal end of the catheter.

Preferably, the blood pump can be implanted as a rotary blood pump, i.e. a blood pump driven by a rotary motor.

The blood pump may be catheter-based for implantation or placement directly into the heart via a corresponding blood vessel. For example, the blood pump may be a blood pump as disclosed in, for example, US 5 911 685, which is arranged for temporary placement or implantation into the left or right heart of a patient. As mentioned above, the present invention is particularly useful for intravascular blood pumps and less relevant for larger VADs which are not placed intravascularly or intra-cardiacly but are implanted outside the heart of a patient, for example in the thoracic cavity.

Preferably, the blood pump is a low-inertia device. (a) the blood pump is a low-inertia device by having one or more of the following features: (b) the moving parts of the blood pump, in particular the rotating parts, e.g. the rotor or the impeller, are made of a low-weight material, e.g. plastic, and thus have a low mass; (c) the drive unit, e.g. an electric motor, is arranged close to, preferably very close to, most preferably adjacent to, the moving parts, e.g. the rotor or the impeller, driven by the drive unit; (d) if the blood pump is catheter-based, there is no rotating drive cable or drive wire; (e) the rotating parts of the pump unit, e.g. the rotor or the impeller, driven by the drive unit, and the connection or joint, e.g. the shaft, of the drive unit are short; and (f) all moving parts of the blood pump, in particular the rotating parts, have a small diameter.

Low-inertia devices include intravascular blood pumps, in particular for percutaneous insertion into a patient's blood vessels. Due to their small diameter (especially compared to relatively bulky VADs), all moving parts of the intravascular blood pump are light in weight and are located close to the axis of rotation. This allows for a very precise control of the pump speed, since the rotation of the impeller is only slightly influenced by the inertia of the impeller. This means that only a slight delay occurs between the command signal and the actual response of the blood pump. In contrast, VADs, which are configured as centrifugal blood pumps for example, can have large rotors with large diameters and therefore larger masses, and therefore cannot be referred to as "low-inertia devices".

One feature of low-inertia devices is that, for example, reducing the pump speed of a low-inertia device, especially reducing the pump speed quickly, does not require a negative speed signal (which is typically needed in large VADs) or another brake command, and reducing the motor current directly results in a reduced pump speed, and the blood pump can enter zero-flow control mode simply by reducing the motor current. This is particularly relevant for intrabeat control, since the cardiac cycle is very short and requires a short reaction time of the moving part of the blood pump. Conversely, it is equally desirable to rapidly accelerate the moving part to exit zero-flow control mode, i.e. rapidly increase the pump speed.

For example, with a low-inertia device in the sense of the present invention, the pump speed can be significantly increased or decreased within a very short period of time, for example within about 50 ms to about 100 ms, preferably within 60 ms to 80 ms. In other words, the pump speed can be changed rapidly or in a "step change".

For example, within the above time period, the pump speed may be decreased from about 35,000 rpm to about 10,000 rpm, or in another blood pump, from about 51,000 rpm to about 25,000 rpm, or conversely increased. However, the duration of the pump speed change also depends on other factors, such as blood flow, pressure differential, the intended amount of the pump speed change (i.e., the difference between the pump speed before and after the intended speed change), or the timing within the cardiac cycle (because blood flow accelerates during systole and slows down during diastole).

A third aspect provides a method for controlling a blood flow (Q _pump (t)) of an intravascular blood pump as discussed together with the first aspect. Namely, the blood pump comprises a pump unit having a drive unit and is configured to deliver blood from a blood flow inlet toward a blood flow outlet. The method comprises the steps of: (i) comparing a set blood flow value (Q _pump ^set (t)) with a blood flow value (Q _pump (t)) to obtain a control error (e(t)) in a first closed-loop cycle; (ii) determining a set speed value (n _pump ^set (t)) for the drive means from the control error (e(t)); and controlling a drive speed (n _pump (t)) of the drive unit by comparing the set speed value (n _pump ^set (t)) with the drive speed (n _pump (t)) in a second closed-loop cycle.

The method preferably further comprises the step of providing a zero flow mode in which the set blood flow value (Q _pump ^set (t)) is zero for a preset zero flow control period, and preferably the step of setting the preset zero flow control period to last for a portion of one cardiac cycle of the assisting heart, or for at least one complete cardiac cycle or a preset number of consecutive cardiac cycle fractions and/or complete cardiac cycles.

As described in more detail above with respect to the control device, the method may include a step of determining or estimating blood flow (Q _pump (t)) using a suitable lookup table.

Preferably, the first closed-loop cycle is an outer control loop and the second closed-loop cycle is an inner control loop of the cascade control. A cascade control system comprising an outer control loop and an inner control loop has been described in more detail above with respect to the control device and is also valid with respect to the method.

Preferably, the method further comprises the step of synchronizing the zero flow control period with at least one specific characteristic cardiac cycle event.

Preferably, the start and/or end of the zero flow control period is synchronized with the occurrence of at least one characteristic cardiac cycle event.

Preferably, the method further comprises the step of monitoring one or more values of characteristic cardiac parameters.

Preferably, the method further comprises the step of identifying a trend in one or more monitored values of the characteristic cardiac parameter.

A fourth aspect provides a control device according to the first aspect configured to perform a method according to the third aspect.

The above-described functions and functionality of the control device and the corresponding control method can be implemented by a corresponding computational unit of the control device, either in hardware or in software or in a combination thereof. This computational unit can be configured by a corresponding computer program having software codes for causing the computational unit to perform each required control step. Such programmable computational units are fundamentally well known in the art and to the person skilled in the art. Therefore, it is not necessary to describe such programmable computational units in detail here. Furthermore, the computational unit can comprise specific designated hardware useful for specific functions, such as, for example, one or more signal processors for processing and/or analyzing the measured signals discussed. Furthermore, each unit for controlling the speed of the drive unit of the blood pump can also be implemented by a respective software module.

The corresponding computer program may be stored on a data carrier containing the computer program. Alternatively, the computer program may be transmitted in the form of a data stream containing the computer program, for example via the Internet, without the need for a data carrier.

Hereinafter, the present invention will be described by way of example with reference to the accompanying drawings.
Figure 1 shows a block diagram of feedback control.
Figure 2 illustrates an exemplary blood pump extending through the aortic valve into the left ventricle and positioned through the aorta, together with a block diagram of a control device for the pumping speed of the blood pump.
Figure 3 illustrates the blood pump of the example of Figure 1 in more detail.
FIG. 4 is an exemplary diagram illustrating a set of characteristic curves representing the relationship between the actual pressure difference between the inlet and outlet of the blood pump, the actual pumping speed of the blood pump, and the blood flow generated through the blood pump.

FIG. 1 is a block diagram illustrating an example of a feedback control loop for blood flow control implemented as a cascade control system. The control loop includes an outer controller (401) and an inner controller (402). The outer controller (401) is embedded in the outer control loop and regulates the generated blood flow (Q _pump (t)) by comparing a blood flow command signal (Q pump ^set (t)) with a generated blood flow (Q _pump (t)) of the blood pump (50), for example, as illustrated in FIG. 3, and by setting a set-point of the inner control loop, i.e., a speed command signal (n _pump ^set (t)) of the blood _pump (50). The inner controller (402) is part of the inner control loop and controls the speed (n _pump (t)) of the blood pump (50) by adjusting a motor current (I _pump (t)) accordingly.

In the feedback loop illustrated in FIG. 1, the generated blood flow (Q _pump (t)) is exemplarily calculated by a lookup table representing the relationship between, for example, the current (I _pump (t)), the velocity (n _pump (t)), and the generated blood flow (Q _pump (t)). Alternatively or additionally, another lookup table can be used to represent the relationship between the pressure difference between the blood pump outlet (56) and the blood pump inlet (54) (see FIG. 3), the velocity (n _pump (t)), and the blood flow (Q _pump (t)) _. Another alternative or additional option for acquiring data of the generated blood flow (Q pump (t)) is to use a flow sensor.

The flow control regulates the blood flow (Q _pump (t)) through the blood pump (50) according to a blood flow command signal (Q _pump ^set (t)) which can be a constant value (also called set-value) or a signal which varies over time. The constant blood flow set-value (Q _pump ^set (t)) is in the range of [-5 ... 10] L/min, preferably in the range of [0 ... 5] L/min, and most preferably can be a very low blood flow such as 0 L/min or zero flow.

One of the purposes of the flow control disclosed herein is to monitor values of characteristic parameters of the heart with an implanted pump (50) to determine the state of recovery of the heart while reducing the impact of the pump on cardiac function. For this purpose, the flow control can use a set blood flow of 0 L/min (Q _pump ^set (t)) or a very low blood flow, such as zero flow.

It was found that the inner control loop can have a much smaller time constant than the outer control loop. As such, the inner control loop responds much faster than the outer control loop. In addition, the inner control loop can be performed at a higher sampling rate than the outer control loop.

For example, the sampling rate of data in the inner control loop (fs _IN ) can be in the range of [250...10k] Hz, preferably in the range of [1...3] kHz, and most preferably 2.5 kHz.

For example, the sampling rate of data in the outer control loop (fs _OUT ) can be in the range of [25…1000] Hz, preferably in the range of [100…300] Hz, and most preferably 250 Hz.

Figures 2 and 3 illustrate an example of a blood pump. The blood pump is an intravascular blood pump configured for percutaneous insertion into the heart. In the illustrated embodiment, the blood pump is a micro-axial rotary blood pump (hereinafter briefly referred to as blood pump (50)). Such a blood pump is known, for example, from US 5 911 685 A.

The blood pump (50) is based on a catheter (20), by means of which the blood pump (50) can be temporarily introduced via a blood vessel into a ventricle of a patient's heart. The blood pump (50) comprises, in addition to the catheter (20), a rotary drive unit (51) fixed to the catheter (20). The rotary drive unit (51) is coupled to a pump unit (52) located at some axial distance therefrom.

The flow cannula (53) is connected to the pump unit (52) at one end thereof, extends from the pump unit (52) and has a blood flow inlet (54) located at its other end. The blood flow inlet (54) has a soft flexible tip (55) attached thereto.

The pump unit (52) includes a pump housing having a blood flow outlet (56). In addition, the pump unit (52) includes a drive shaft (57) protruding from the drive unit (51) into the pump housing of the pump unit (52). The drive shaft (57) drives an impeller (58) as a thrust element. During operation of the blood pump (50), blood is sucked in through the blood flow inlet (54), delivered through the cannula (53), and discharged through the blood flow outlet (56). The blood flow is generated by the rotating impeller (58) driven by the drive unit (51).

In the illustrated embodiment, three lines pass through the catheter (20), namely two signal lines (28A, 28B) and a power supply line (29) for powering the drive unit (51) of the blood pump (50). The signal lines (28A, 28B) and the power supply line (29) are attached at their proximal ends to a control device (100) (Fig. 2). The signal lines (28A, 28B) are associated with respective blood pressure sensors having corresponding sensor heads (30 and 60) respectively. The power supply line (29) comprises a supply line for powering the drive unit (51).

The drive unit (51) may be a synchronous motor. In an exemplary configuration, the electric motor may include several motor winding units for driving an impeller (58) coupled with a drive shaft (57). The rotor of the synchronous motor may include at least one field winding, or alternatively, permanent magnets in the case of a permanent magnet excited synchronous motor.

In a preferred embodiment, the blood pump (50) is a catheter-based micro-axial rotary blood pump for percutaneous insertion into the heart of a patient via a blood vessel of the patient. Here, "micro" indicates that the blood pump is small enough in size to be inserted percutaneously into the heart via a blood vessel leading to the heart, for example into one of the ventricles of the heart. This also defines the blood pump (50) as an "intravascular" blood pump for percutaneous insertion. Here, "axial" indicates that the pump unit (52) and the drive unit (51) driving it are arranged in an axial configuration. Here, "rotary" means that the functionality of the pump is based on the rotational motion of the impeller (58) driven by the thrust element, i.e. the rotary electric motor of the drive unit (51).

As discussed above, the blood pump (50) is based on a catheter (20), by means of which insertion of the blood pump (50) through a blood vessel can be performed, and through which a power supply line (29) can be passed to supply power to the drive unit (51) and, for example, to supply control signals from the drive unit (51) and the sensor head (30, 60).

As mentioned above, the present invention is particularly designed for intravascular blood pumps, for example blood pumps (50) as illustrated in FIG. 3, and is rarely or even unsuitable for VADs which are implanted outside the heart of a patient, for example centrifugal blood pumps which are positioned in the thoracic cavity and connected to the heart of the patient and which operate at different pump speeds. As explained herein, this is due to inertial effects which significantly affect the function of large VADs but can be avoided in low inertia devices such as intravascular blood pumps.

As illustrated in FIG. 2, each signal line (28A, 28B) is connected to one respective blood pressure sensor having a corresponding sensor head (30 and 60), which heads are positioned externally on the housing of the pump unit (52). The sensor head (60) of the first pressure sensor is connected to the signal line (28B) and is intended to measure the blood pressure at the blood flow outlet (56). The sensor head (30) of the second blood pressure sensor is connected to the signal line (28A) and is intended to measure the blood pressure at the blood flow inlet (54). Basically, the signals captured by the pressure sensors (carrying the respective information about the pressure at the location of the sensor and which can be of any suitable physical origin, for example optical, hydraulic or electrical origin) are transmitted via the respective signal lines (28A, 28B) to the corresponding input of the data processing unit (110) of the control device (100). In the example illustrated in FIG. 2, the blood pump (50) is positioned in the aorta via the aortic valve in the left ventricle of the heart, and the pressure sensor is arranged to measure the aortic pressure (AoP(t)) by the sensor head (60) and the ventricular pressure (LVP(t)) by the sensor head (30).

The data processing unit (110) is configured for acquisition of the outer and inner signals for signal processing, including for example computation of the difference between the pressure signals as a basis for estimating the generated blood flow (Q _pump (t)) which can serve as a control signal for a flow control approach, signal analysis for detecting the occurrence of a characteristic event during the cardiac cycle based on the acquired and computed signals, and for generating a trigger signal (σ(t)) for triggering a velocity command signal generator (120). For the given example of the flow control approach, the velocity command signal generator (120) represents the outer controller (401) of FIG. 1.

In the illustrated embodiment, the data processing unit (110) is connected via signal lines corresponding to additional measurement devices, which are illustrated generally by 300. These additional measurement devices are, in the present embodiment, a patient monitoring unit (310) and an electrocardiograph (ECG; 320); obviously, these two devices (310 and 320) are only two examples and are not exhaustive; that is, other measurement devices may also be used to provide useful signals. The illustrated ECG (320) provides an ECG signal (ECG(t)) to the data processing unit (110).

The control device (100) further includes a user interface (200). The user interface (200) is for interaction with a user of the device. The user interface (200) includes a display (210) as an output means and a communication interface (220) as an input means. On the display (210), values of set parameters, values of monitored parameters, for example, measured pressure signals, and other information are displayed. In addition, by means of the communication interface (220), the user of the control device (100) can control the control device (100) by changing the setup and settings of the entire system, for example, consisting of the blood pump and the control device (100).

For a given example of a flow control approach, one setting could be the selection of a desired pump flow (Q _pump ^set (t)) in Fig. 1.

The data processing unit (110) is configured to derive or predict the occurrence time of one or more predefined characteristic events during a cardiac cycle of the assisted heart. For example, the data processing unit (110) is configured to detect predefined characteristic cardiac cycle events during a cardiac cycle by real-time analysis of the monitored signal. Alternatively or additionally, the predefined characteristic cardiac cycle events, for example, R-waves, may be identified by an ECG signal from the ECG (320).

The occurrence of one or more determined predefined characteristic events is used for the generation of a specific trigger signal (σ(t)) or a sequence of trigger signals (σ(t)). The resulting trigger signal (σ(t)) (or a sequence thereof) is sent to a speed command signal generator (120) to correspondingly trigger a speed command signal change provided to a speed control unit (130).

In the context of the present invention, the speed command signal generator (120) is configured to operate the blood pump (50) in zero flow control mode.

The data processing unit (110) may be configured to predict the time of occurrence of at least one predefined characteristic cardiac cycle event in an upcoming cardiac cycle based on stored information about characteristic cardiac cycle events occurring during the current and/or previous cardiac cycles, and also to analyze previous values of such rate command signal (n _pump ^set (t)).

For example, a characteristic cardiac cycle event may be the onset of a contraction of the heart at the beginning of systole. The detected or predicted occurrence of such a characteristic cardiac cycle event may be used to synchronize the sequential application of a specific control approach to the blood pump (50) within a specific time interval of the cardiac cycle, or within one or several cardiac cycles.

Correspondingly, the speed command signal generator (120) is configured to adjust a speed command signal (n _pump ^set (t)) to the blood pump (50) to control the generated blood flow (Q _pump (t)) according to a given blood flow command signal (Q _pump ^set (t)), which may be set to, for example, 0 L/min.

In order to control the generated blood flow (Q _pump (t)), the speed command signal generator (120) is configured as an outer controller of the cascade control system to provide a suitable speed command signal (n _pump set (t)) to the speed control unit (130) in a time-continuous way (as a first setup) by continuously controlling the generated blood flow (Q _pump ⁽ t)) or in an event-based switching control way (as a second setup).

In the first setup, the command signal generator (120) is part of a cascade blood-flow control system that is supplied with external and internal signals by the data processing unit (110) and continuously provides a speed command signal (n _pump ^set (t)) to the speed control unit (130).

In the second setup, the speed command signal generator (120) operates as in the first setup with the additional feature of switching on and off continuous blood flow control.

In zero flow control mode, the speed command signal (n _pump ^set (t)) is continuously regulated by the flow controller in the outer control loop. On/off switching is triggered by at least one trigger signal (σ(t)) provided by the data processing unit (110).

The second setup is suitable if the zero flow control is applied only for a short time interval, in particular short compared to the duration of one cardiac cycle; in other words, the generated blood flow is controlled with a blood flow command signal of 0 L/min (Q _pump ^set (t)) only for a short time interval within the cardiac cycle (blood flow control within one beat).

The speed control unit (130) controls the speed (n pump (t)) of the blood pump (50) according to a speed command signal (n _pump ^set (t)) by providing current (I _pump (t)) to the drive unit (51) of the blood pump (50) through the power _{supply line} (29).

The current level of the supplied motor current (I _pump (t)) corresponds to the current currently required ^by the drive unit (51), for example by _the electric motor, to achieve a target speed level as defined by the speed command signal (n pump set (t)). A measurement signal such as the supplied motor current (I _pump (t)) can be used as a representative signal of the internal signal of the control device (100) and can be provided to a data processing unit (110) for further processing. Via the power supply line (29), the blood pump (50) can also communicate with the control unit (100).

Basically, first of all, the control device (100) is configured to operate the blood pump (50) in a selectable zero flow control mode, in which the blood flow (Q _pump (t)) of the blood pump (50) is controlled to respond to a changing pressure difference between the blood outlet (56) and the blood inlet (54) due to a heartbeat, which may be considered as a disturbance. The blood flow (Q _pump (t)) is controlled by modulating a speed command signal (n _pump ^set (t)). As proposed herein, the control device (100) is configured to control the blood flow (Q _pump (t)) of the blood pump (50) such that the blood pump (50) generates zero blood flow during a preset zero flow control period.

In a first setup having continuous flow control, the preset zero flow control period is set to last for at least one complete cardiac cycle or a preset number of consecutive complete cardiac cycles. Also in the first setup, the control device (100) is configured to monitor values of one or more characteristic cardiac parameters with the implanted blood pump (50). Furthermore, the monitored values of the one or more characteristic cardiac parameters can be used as an indicator of the state of cardiac recovery.

In a second setup with event-based zero flow control, the preset zero flow control period is set to be a portion of the duration of one cardiac cycle of the heart with the implanted blood pump (50). In this setup, the control device (100) is configured to synchronize the occurrence of certain characteristic cardiac cycle events with the start and end of the zero flow control period.

In particular, the control device (100) can control the blood flow (Q _pump (t)) through the blood pump (50) periodically or randomly.

In a particular implementation, a characteristic cardiac cycle event is opening of the aortic valve or closure of the aortic valve or opening of the mitral valve or closure of the mitral valve, or a particular pressure value that is an end-diastolic left ventricular pressure.

Also in the second setup, as in the first setup, the control device (100) is configured to monitor the values of one or more characteristic cardiac parameters of the heart with the implanted blood pump (50) during the zero flow control period. The monitored values of the one or more characteristic cardiac parameters can also be used as an indicator of the state of cardiac recovery.

The control device (100) is further configured to identify trends in the values of one or more monitored characteristic parameters. As noted above, the trends may also be interpreted as indicators of cardiac recovery status.

In any case, in order to implement the zero flow control mode, the control device (100) is configured to control the blood flow (Q _pump (t)) by adjusting the speed command signal (n _pump ^set (t)) of the blood pump (50), whereby the driving speed (n _pump (t)) is affected by the changing blood pressure difference between the blood flow outlet (56) of the blood pump (50) and the blood flow inlet (54) of the blood pump (50) during the cardiac cycle.

In particular, the control device (100) is configured to determine the blood flow (Q _pump (t)) of the blood pump (50) based on preset signals, for example, the driving speed (n _pump (t)), the current (I _pump (t)), and/or the pressure difference between the blood flow outlet (56) and the blood flow inlet (54) of the blood pump (50).

FIG. 4 is an exemplary diagram showing a set of characteristic curves representing the relationship between a pressure difference (ΔP _pump (t)) between a blood flow outlet (56) and a blood flow inlet (54) of a blood pump (50), a driving speed (n _pump (t)) of the blood pump (50), and a blood flow (Q _pump (t)) generated by the blood pump (50) through, for example, a flow cannula (53) of FIG. 3.

To perform zero flow control, the data processing unit (110) is configured to continuously determine a blood flow (Q _pump (t)) generated by the blood pump (50) based on a known speed (n _pump (t)), a known current (I pump (t)) supplied to the pump unit (51), and/or a monitored pressure difference (ΔP _pump (t)) between a blood flow outlet (56) and a blood flow inlet (54) of the blood _pump (50). The set blood flow value (Q _pump ^set (t)) can be zero or at least a positive value close to zero.

For example, based on FIG. 4, when the monitored pressure difference (ΔP _pump (t)) between the blood flow outlet (56) and the blood flow inlet (54) of the blood pump (50) is 60 mmHg, the driving speed (n _pump (t)) should be about 20.000 1/min (rpm) to generate a blood flow (Q _pump (t)) of about 0 L/min.

It will be appreciated that the values, relationships and shapes of the curves depicted in the characteristic diagram of Fig. 4 are merely examples and may vary depending on the blood pump used, the patient or other factors. In particular, each and every blood pump, even blood pumps of the same type, may have an individual characteristic diagram. That is, the lookup table is pump specific. Furthermore, once implanted in a patient, the characteristic diagram may need to be adjusted by patient-specific correction factors that include various factors such as blood viscosity, position of the blood pump, etc. Using correction factors can increase the accuracy of the flow estimates obtained from the lookup table.

As discussed above, the current pressure difference (ΔP _pump (t)) can be determined by a pressure sensor of the blood pump (50) (e.g., sensor (30, 60, FIG. 3)). Accordingly, the speed control unit (130) can be continuously provided with values from a storage unit, for example, a look-up table in which the characteristic curves of FIG. 4 are stored (wherein representing the relationship between the values (ΔP _pump (t), Q _pump (t) and n _pump (t)) discussed above). The storage unit can be a read only memory of the data processing unit (110) or, alternatively, a storage chip within the blood pump (50) or within the control console (130) of the data processing unit.

At least one characteristic cardiac parameter value is at least one of the arterial blood pressures measured whenever a zero flow mode of operation is established.

Preferably, the blood pump (50) is a low-inertia device. This is achieved in particular by the fact that the moving, in particular rotating parts of the blood pump (50), for example the rotor or the impeller, are made of low-weight materials, for example plastic, and thus comprise a low mass. Additionally, the drive unit, for example an electric motor, is arranged close to, preferably very close to, most preferably adjacent to, a thrust element driven by the drive unit, for example the rotor or the impeller (58). Additionally, even though the blood pump (50) is catheter-based, there are no rotating drive cables or drive wires. Additionally, the connection or joint between the thrust element driven by the drive unit (51), for example the rotor or the impeller (58), and the drive unit (51), for example the shaft (57), is kept short. Additionally, all moving, in particular rotating parts of the blood pump (50) have a small diameter.

In summary, in the zero flow control approach proposed herein, the control device (100) controls the generated blood flow (Q _pump (t)) through the blood pump (50) by a cascade control consisting of an outer control loop and an inner control loop. This means that the generated blood flow (Q _pump (t)) through the blood pump (50) is controlled by regulating the speed command signal (n _pump ^set (t)) to the drive unit (51) of the blood pump (50) in the outer control loop, and the drive speed (n _pump (t)) is controlled in the inner control loop by regulating the current (I _pump (t)). The zero flow control approach is applied continuously or partially continuously, i.e. the zero flow control period lasts for one or several complete cardiac cycles or only a part of a cardiac cycle. In case the preset zero flow control period lasts only a part of the cardiac cycle duration, the zero flow control period can be synchronized with the heartbeat by at least one characteristic event of the cardiac cycle.

Claims (25)

A control device (100) for controlling blood flow (Q _pump (t)) of an intravascular blood pump (50) for percutaneous insertion into a patient's blood vessel,
The above blood pump (50) includes a pump unit (52); and a driving unit (51) for driving the pump unit configured to deliver blood from a blood flow inlet (54) toward a blood flow outlet (56);
The control device (100) is configured to operate the blood pump (50) in a selectable zero flow control mode - a blood flow command signal (Q _pump ^set (t)) is selected - and the control device comprises a first controller (401) and a second controller (402),
A control device, wherein the first controller (401) is configured to control the blood flow (Q _pump (t)) by regulating a speed command signal (n _pump ^set (t)) for the driving unit (51), and the second controller (402) is configured to control the driving speed (n _pump (t)) of the driving unit (51). In the first paragraph,
The control device (100) wherein the first controller (401) is further configured to determine the speed command signal (n _pump ^set (t)) based on the difference (ΔQ) between the blood flow command signal (Q _pump ^set (t) ₎ and the blood flow (Q pump (t)). In paragraph 1 or 2,
The second controller (402) is a control device (100) configured to control the driving speed (n _pump (t)) by adjusting the driving current (I _pump (t)) supplied to the driving unit (51). In any one of claims 1 to 3,
A control device (100), wherein the first controller (401) and the second controller (402) are parts of a cascade control system, and in the cascade system, the first controller (401) is an outer controller and the second controller is an inner controller (402). In any one of claims 1 to 4,
The control device (100) is configured to control the blood flow (Q _pump (t)) during a preset zero flow control period. In paragraph 5,
The above preset zero flow control period is set to last for a portion of one cardiac cycle of the assisted heart; or
A control device (100) wherein the preset zero flow control period is set to last for at least one complete cardiac cycle or a preset number of consecutive complete cardiac cycles. In any one of claims 1 to 6,
The control device (100) is configured to synchronize the occurrence of at least one characteristic cardiac cycle event with the zero flow control period. In Article 7,
A control device (100) wherein the start and/or end of said zero flow control period is synchronized with the occurrence of said at least one characteristic cardiac cycle event. In Article 8,
A control device (100) wherein at least one of the characteristic cardiac cycle events is opening of the aortic valve or closing of the aortic valve. In any one of claims 1 to 9,
A control device (100) configured to monitor the values of one or more characteristic cardiac parameters. In any one of claims 1 to 10,
The control device (100) is configured to periodically or arbitrarily operate the blood pump (50) in the zero flow control mode. In any one of claims 1 to 11,
A control device (100) configured to identify a trend in one or more values of a monitored characteristic cardiac parameter. In Article 12,
A control device (100) wherein the at least one characteristic cardiac parameter is at least one of arterial pressure pulsatility (AOP| _max - AOP| _min ), mean arterial pressure, cardiac contractility (dLVP(t)/dt| _max ), cardiac relaxation (dLVP(t)/dt| _min ), and heart rate (HR). In any one of claims 1 to 13,
The control device (100) is configured to measure the blood flow (Q _pump (t)) by a sensor or to calculate or estimate the blood flow (Q _pump (t)). In any one of claims 1 to 14,
The control device (100) is configured to determine the blood flow (Q _pump (t)) using a look-up table, wherein the look-up table represents a relationship between the blood flow (Q _pump (t)) and the driving speed (n _pump (t)), at least one of a pressure difference (ΔP _pump (t)) between the blood flow outlet (56) and the blood flow inlet (54), and a driving current (I _pump (t)) supplied to the driving unit (51). An intravascular blood pump (50) for percutaneous insertion into a patient's blood vessel; and
A control device (100) comprising any one of claims 1 to 15,
System. In Article 16,
The above blood pump (50) is a low inertia device including at least one of the following features:
The following features are:
The moving parts of the blood pump, particularly the rotating parts, for example the rotor or the impeller, are made of a low-weight material, for example plastic, and thus have a low mass;
The drive unit (51), for example an electric motor, is arranged close to, preferably very close to, most preferably adjacent to, a moving part driven by the drive unit (51), for example a rotor or an impeller, and if catheter-based, preferably without a rotating drive cable or drive wire;
A rotating part driven by the drive unit (51), for example, a joint or connecting part of the rotor or impeller and the drive unit (51), for example, a point where the shaft is short; and
A system in which all moving parts of the above blood pump, particularly the rotating parts, have small diameters. A method for controlling blood flow (Q _pump (t)) of an intravascular blood pump (50) for percutaneous insertion into a patient's blood vessel,
The above blood pump (50) includes a pump unit (52) having a driving unit (51) and is configured to deliver blood from a blood flow inlet (54) toward a blood flow outlet (56).
The above method,
A step of obtaining a control error (e(t)) in a first closed-loop cycle by comparing the set blood flow value (Q _pump ^set (t)) and the blood flow value (Q _pump (t));
A step of determining a set speed value (n _pump ^set (t)) for the driving means from the above control error (e(t)); and
A step of controlling the driving speed (n _pump (t)) of the driving unit (51) by comparing the set speed value (n _pump ^set (t)) and the driving speed (n _pump (t)) in the second closed-loop cycle,
method. In Article 18,
A step of providing a zero flow mode in which the set blood flow value (Q _pump ^set (t)) is zero during a preset zero flow control period; and
The method preferably further comprising the step of setting the preset zero flow control period to last for a portion of one cardiac cycle of the assisted heart, or for at least one complete cardiac cycle or a preset number of consecutive cardiac cycle fractions and/or complete cardiac cycles. In Article 18 or 19,
A method wherein the first closed-loop cycle is an outer control loop and the second closed-loop cycle is an inner control loop of cascade control. In any one of Articles 18 to 20,
A method further comprising the step of synchronizing said zero flow control period with said at least one specific characteristic cardiac cycle event. In any one of Articles 18 to 21,
A method wherein the start and/or end of said zero flow control period is synchronized with the occurrence of said at least one characteristic cardiac cycle event. In any one of Articles 18 to 22,
A method further comprising the step of monitoring one or more values of characteristic cardiac parameters. In any one of Articles 18 to 23,
A method further comprising the step of identifying a trend in said one or more monitored values of said characteristic cardiac parameters. A control device (100) of any one of claims 1 to 14, configured to execute a method of any one of claims 18 to 24.